Field of the Invention
The present invention generally relates to systems and methods of obtaining images and spectra of anatomical structures using magnetic resonance (MR) systems. More particularly, the present invention pertains to multiple embodiments of a multichannel surface coil array and associated interface devices capable of providing images and spectroscopic results from the MR signals obtained from the nuclei excited during MR procedures.
Description of Related Art
The following background information is provided to assist the reader to understand the invention disclosed below and the environment in which it will typically be used. The terms used herein are not intended to be limited to any particular narrow interpretation unless clearly stated otherwise, either expressly or implied, in this document.
Magnetic resonance imaging (MRI) is a noninvasive method of producing high quality images of the interior of the human body. It allows medical personnel to see inside the human body without surgery or the use of ionizing radiation such as X-rays. The images are of such high resolution that cancer and other forms of pathology can often be visually distinguished from healthy tissue. Magnetic resonance techniques and systems have also been developed for performing spectroscopic analyses by which the chemical content of body tissue or other material can be ascertained.
MRI uses a powerful magnet, radio waves, and computer technology to create detailed images of the soft tissues, muscles, nerves, and bones in the body. It does so by taking advantage of a basic property of the hydrogen atom, an atom found in abundance in all cells within living organisms. In the absence of a magnetic field, the nuclei of hydrogen atoms spin like a top, or precess, randomly in every direction. When subject to a strong magnetic field, however, the spin-axes of the hydrogen nuclei align themselves in the direction of the field. This is because the nucleus of the hydrogen atom has what is referred to as a large magnetic moment, a strong inherent tendency to line up with the direction of the field. Collectively, the hydrogen nuclei of the area to be imaged create an average vector of magnetization that points parallel to the magnetic field.
A typical MRI system, or scanner, includes a main magnet, three gradient coils, a radio frequency (RF) antenna (often referred to as the whole body coil), and a computer station from which an operator can control the system. The chief component of the MRI system, however, is the main magnet. It is typically superconducting in nature and cylindrical in shape. Within its bore (an opening into which patients are placed during an MRI procedure), the main magnet generates a strong magnetic field, often referred to as the B0 field, which is both uniform and static (non-varying). This B0 magnetic field is oriented along the longitudinal axis of the bore, referred to as the z direction, which compels the magnetization vectors of the hydrogen nuclei in the body to align themselves parallel to that axis. In this alignment, the nuclei are prepared to receive RF energy of the appropriate frequency from the whole body coil. This frequency is known as the Larmor frequency and is governed by the equation ω=γ B0, where ω is the Larmor frequency (at which the hydrogen atoms precess), γ is the gyromagnetic constant, and B0 is the strength of the static magnetic field.
The RF antenna, or whole body coil, is generally used both to transmit pulses of RF energy and to receive the resulting MR signals induced thereby in the hydrogen nuclei. Specifically, during its transmit cycle, the body coil broadcasts RF energy into the cylindrical bore. This RF energy creates a radio frequency magnetic field, also known as the RF B1 field, whose magnetic field lines are directed in a line perpendicular to the magnetization vector of the hydrogen nuclei. The RF pulse causes the spin-axes of the hydrogen nuclei to tilt with respect to the main (B0) magnetic field, thus causing the net magnetization vector to deviate from the z direction by a known angle. The RF pulse, however, will affect only those hydrogen nuclei that are precessing about their axes at the frequency of the RF pulse. In other words, only the nuclei that “resonate” at that frequency will be affected, and such resonance is achieved in conjunction with the operation of the three gradient coils.
The gradient coils are electromagnetic coils. Each gradient coil is used to generate a linearly varying yet static magnetic field along one of the three spatial directions (x, y, z) within the cylindrical bore known as the gradient B1 field. Positioned inside the main magnet, the gradient coils are able to alter the main magnetic field on a very local level when they are turned on and off very rapidly in a specific manner. Thus, in conjunction with the main magnet, the gradient coils can be operated according to various imaging techniques so that the hydrogen nuclei, at any given point or in any given strip, slice, or unit of volume, will be able to achieve resonance when an RF pulse of the appropriate frequency is applied. In response to the RF pulse, the precessing hydrogen atoms in the selected region absorb the RF energy being transmitted from the body coil, thus forcing the magnetization vectors thereof to tilt away from the direction of the main (B0) magnetic field. When the body coil is turned off, the hydrogen nuclei begin to release the RF energy in the form of the MR signal, as explained further below.
One well known technique that can be used to obtain images is referred to as the spin echo imaging technique. Operating according to this technique, the MRI system first activates one gradient coil to set up a magnetic field gradient along the z-axis. This is called the “slice select gradient”, and it is set up when the RF pulse is applied and it shuts off when the RF pulse is turned off. It allows resonance to occur only within those hydrogen nuclei located within a slice of the area being imaged. No resonance will occur in any tissue located on either side of the plane of interest. Immediately after the RF pulse ceases, all of the nuclei in the activated slice are “in phase”, i.e., their magnetization vectors all point in the same direction. Left to their own devices, the net magnetization vectors of all the hydrogen nuclei in the slice would relax, thus realigning with the z direction. Instead, however, the second gradient coil is briefly activated to create a magnetic field gradient along the y-axis. This is called the “phase encoding gradient”. It causes the magnetization vectors of the nuclei within the slice to point, as one moves between the weakest and strongest ends of the gradient, in increasingly different directions. Next, after the RF pulse, slice select gradient, and phase encoding gradient have been turned off, the third gradient coil is briefly activated to create a gradient along the x-axis. This is called the “frequency encoding gradient” or “read out gradient”, as it is only applied when the MR signal is ultimately measured. It causes the relaxing magnetization vectors to be differentially re-excited, so that the nuclei near the low end of the gradient begin to precess at a faster rate, and those at the high end pick up even more speed. When these nuclei relax again, the fastest ones (those which were at the high end of the gradient) will emit the highest frequency of radio waves.
Collectively, the gradient coils allow the MR signal to be spatially encoded, so that each portion of the area being imaged is uniquely defined by the frequency and phase of its resonance signal. In particular, as the hydrogen nuclei relax, each becomes a miniature radio transmitter giving out a characteristic pulse that changes over time, depending on the local microenvironment in which it resides. For example, hydrogen nuclei in fats have a different microenvironment than do those in water, and thus transmit different pulses. Due to these differences, in conjunction with the different water-to-fat ratios of dissimilar tissues, different tissues transmit radio signals of different frequencies. During its receive cycle, the body coil detects these miniature radio transmissions, which are often collectively referred to as the MR signal. From the body coil, these unique resonance signals are conveyed to the receivers of the MR system, where they are converted into mathematical data corresponding thereto. The entire procedure must be repeated multiple times to form an image with a good signal-to-noise ratio (SNR). Using multidimensional Fourier transformations, an MR system can convert the mathematical data into a two- or even a three-dimensional image.
When more detailed images of a specific part of the body are needed, a local coil is often used instead of the whole body coil. A local coil can take the form of a volume coil or a surface coil. A volume coil is used to surround or enclose the volume to be imaged (e.g., a head, an arm, a wrist, a leg, or a knee). A surface coil, however, is merely placed upon the surface of a patient so that the underlying region of interest (e.g., the abdominal, thoracic, and/or pelvic regions) can be imaged. In addition, a local coil can be designed to operate either as a receive-only coil or a transmit/receive (T/R) coil. The former is only capable of detecting the MR signals produced by the body in response to an MRI procedure, as noted above. A T/R coil, however, is capable of both receiving the MR signals as well as transmitting the RF pulses that produce the RF B1 magnetic field, which is the prerequisite for inducing resonance in body tissue.
It is well known in the field of MRI to use a single local coil, whether surface or volume, to detect the MR signals. According to the single coil approach, a relatively large local coil is used to cover or enclose the entire region of interest. Early receiving coils were just linear coils, meaning that they could detect only one of the two (i.e., vertical MX′ and horizontal MY′) quadrature components of the MR signals produced by the region of interest. One example of a linear coil is the single loop coil shown in FIG. 1A. This loop is only capable of detecting magnetic fields (i.e., MR signals) that are oriented perpendicular/vertical to the plane of the loop as shown in FIG. 1B. Another example of a linear coil is the butterfly or saddle coil shown in FIG. 2A. Unlike the single loop, the butterfly coil is only sensitive to magnetic fields that are oriented parallel to the plane of the coil as shown in FIG. 2B. This is because a butterfly coil is constructed by twisting a loop in the middle to form two identical subloops about a midpoint. Because the currents flowing in the subloops are the same but flow in counter-rotating directions, the magnetic flux generated by the current flowing through one subloop of the symmetric structure is equal but opposite to the flux due to the current in the other subloop. Therefore, about the midpoint of the structure, the vertical fields, due to the counter-rotating currents, oppose and thus cancel each other. The horizontal fields generated by those currents, however, combine, yielding a magnetic field that is oriented parallel to the plane of the coil.
Accordingly, receiving coils employing quadrature mode detection, meaning that they could intercept both the vertical and horizontal components, have been developed. Compared to linear receiving coils, quadrature receiving coils enabled MRI systems to provide images for which the SNR was much improved, typically by as much as 41%. Even with the improvement brought with quadrature mode detection, the single coil approach still provided images whose quality invited improvement. The disadvantage inherent to the single coil approach is attributable to just one coil structure being used to acquire the MR signals over the entire region of interest.
Phased array coils were also developed to overcome the shortcomings with the single coil approach. Instead of one large local coil, the phased array approach uses a plurality of smaller local coils, with each such coil covering or enclosing only a portion of the region of interest. In a system having two such coils, for example, each of the coils would cover or enclose approximately half of the region of interest, with the two coils typically being partially overlapped for purposes of magnetic isolation. The two coils would acquire the MR signals from their respective portions simultaneously, and they would not interact adversely due to the overlap. Because each coil covers only half of the region of interest, each such coil is able to receive the MR signals at a higher SNR ratio for that portion of the region of the interest within its coverage area. The smaller local coils of the phased array thus collectively provide the MRI system with the signal data necessary to generate an image of the entire region of interest that is higher in resolution than what can be obtained from a single large local coil.
One example of a phased array coil is the Gore® torso array produced by W.L. Gore and Associates, Inc. The torso array contains four surface coils, two of which are disposed in an anterior paddle, and the other two are disposed in a posterior paddle. The two paddles are designed to be placed against the anterior and posterior surfaces, respectively, of the patient about the abdominal, thoracic, and pelvic regions. The torso array is designed for use with an MR system whose data acquisition system has multiple receivers. The four leads of the torso array, one each from the two anterior surface coils and the two posterior surface coils, can be connected to separate receivers, with each receiver amplifying and digitizing the signal it receives. The MR system then combines the digitized data from the separate receivers to form an image whose overall SNR is better than what could be obtained from a single local coil, or even two larger anterior and posterior local coils, covering the entire region of interest alone.
It is also well known to obtain images of internal bodily structures through the use of intracavity probes. An example of a prior art intracavity probe designed primarily for use with 1.0 T and 1.5 T MR systems can be found in U.S. Pat. Nos. 5,476,095 ('095) and 5,355,087 ('087), both of which are assigned to the assignee of the present invention and incorporated herein by reference. The prior art probe disclosed is designed to be inserted into bodily openings, such as the rectum, vagina, and mouth. These patents also disclose interface devices that are designed to interface the prior art intracavity probe with MR imaging and spectroscopy systems. A method of using the intracavity probe is disclosed in U.S. Pat. No. 5,348,010, which is also assigned to the assignee of the present invention and incorporated herein by reference.
The prior art probe, operated in conjunction with its associated interface unit, allows an MR system to generate images of, and spectroscopic results for, various internal bodily structures, such as the prostate gland, colon, or cervix. Examples of such prior art probes include the BPX-15 prostate/endorectal coil (E-coil), the PCC-15 colorectal coil, and the BCR-15 cervix coil, all of which are part of the eCoil™ line of disposable coils produced by MEDRAD, Inc. of Indianola, Pa. Examples of such interface units include the ATD-II and the ATD-Torso units, also produced by MEDRAD, Inc.
The ATD-II unit is used to interface the prior art probe with one receiver of an MR system to provide images or spectra of the region of interest, namely, the prostate gland, colon, or cervix. The ATD-Torso unit is used to interface not only the prior art probe but also the Gore® torso array with multiple receivers of the MR system. When connected to such a probe and the torso array, the ATD-Torso unit allows the MR system to provide images or spectra not only of the prostate gland, colon, or cervix but also of the surrounding anatomy, i.e., the abdominal, thoracic, and pelvic regions.
U.S. Pat. Nos. 7,747,310 and 7,885,704, both of which are assigned to the assignee of the present invention and incorporated herein by reference, disclose several intracavity probes, and associated interface devices, for use with MR systems designed to operate at higher field strengths than the prior art probes of the '087 and '095 patents. For example, the latter reference teaches a probe having a coil loop that includes two drive capacitors and a tuning capacitor, all of which are in series. Connected across each drive capacitor is an output cable having an electrical length of SL+n(λ/4). When each output cable is connected at its other end to the interface device, the coil loop is thereby interconnected through the interface device to the MR system.
With reference to FIG. 3, quadrature intracavity probes have been developed. For instance, International Patent Application Publication No. WO 2010/056911, which is assigned to the assignee of the present invention and is incorporated herein by reference, discloses a single coil structure that is sensitive to both the vertical and horizontal components of the MR signal by virtue of a simple loop-type coil element and a butterfly-type coil element that share a center conductor. More specifically, the quadrature coil, generally designated 10, includes an outer loop 12, a center conductor 14 bisecting the outer loop 12, and an output line, generally designated 16. The outer loop 12 includes a plurality of capacitors including first and second drive capacitors 18 and 20 and first and second tuning capacitors 22 and 24. Of approximately equal values, the drive capacitors 18, 20 are serially deployed within the outer loop 12 and at their junction node 26 form a virtual ground for electrically balancing and impedance matching the loop. Tuning capacitors 22, 24 are also serially deployed within outer loop 12, with their common node 28 being situated diametrically opposite the junction node 26. Of approximately equal values, the tuning capacitors 22, 24 are selected to resonate the outer loop 12 at the operating frequency of the MR system. In that regard, outer loop 12 is shown in FIG. 3 as having four inductors. The values of those inductors merely represent the inductances inherent in the conductive (e.g., copper) segments of the loop. The output line 16 includes two coaxial cables 30 and 32 with the shield conductor of each connected to the junction node 26 of the coil 10. The center conductor 14 extends between and evenly bisects the junction and common nodes 26 and 28 of outer loop 12, and thus maintains the physical and electrical symmetry of quadrature coil 10. FIG. 3 shows the center conductor 14 as having two inductors and a tuning capacitor 34 symmetrically deployed along its length. Like outer loop 12, the values of those inductors merely represent the inductances inherent in the conductor. The value of the tuning capacitor 34 has been selected so that its reactance at the operating frequency equals the inductive reactance of center conductor 14. This configuration permits the simple loop and butterfly elements of the coil to detect MR signals orthogonal and parallel, respectively, to the plane of the coil.
With reference to FIG. 4 and as disclosed in U.S. Pat. No. 7,885,704, a coil having a phased array configuration for use as in an endorectal probe has been developed. The coil includes four coil loops 40, 41, 42, and 43 deployed in a phased array configuration in which each coil loop 40, 41, 42, and 43 is critically overlapped by its neighbor. Each coil loop 40, 41, 42, and 43 includes a drive capacitor 44, 45, 46, and 47 and a tuning capacitor 48, 49, 50, and 51 arranged diametrically opposite to the drive capacitor 44, 45, 46, and 47. In addition, each coil loop 40, 41, 42, and 43 includes an output line 52, 53, 54, and 55 connected across the respective drive capacitor 44, 45, 46, and 47. Accordingly, a four element, four channel configuration is provided. This arrangement provides a demonstrably higher signal-to-noise ratio (SNR) than the quadrature coil 10 described hereinabove with reference to FIG. 3; however, the coverage is less uniform due to the areas of low signal in the critically-coupled (i.e., overlapped conductor) areas. This non-uniformity is undesirable for use in an endorectal probe due to the higher amounts of non-uniformity proximal to the coil conductors.
Despite their widespread acceptance and good reputation in the marketplace, these prior art intracavity probes and interface devices nevertheless have a few shortcomings. For example, they offer limited coverage, exhibit lower signal-to-noise performance, and generally provide less overall flexibility as compared to the endorectal coil technology discussed hereinafter. It is therefore desirable to provide an endorectal coil array and associated interface device capable of providing greater overall flexibility and higher quality images and spectroscopic results from MR signals obtained from nuclei during MR procedures.